Compressed porous materials suitable for implant

ABSTRACT

A high strength porous polymeric material manufactured by a compression process is disclosed. The material results in a network of interconnected collapsed pores, which forces thin overlapping walls and passages to be created. The network provides permeable access for fluid migration throughout the material. The strength and/or permeability are advantageous for medical devices and implants.

CROSS-REFERENCE TO RELATED APPLICATIONS

This patent document is a continuation co-pending U.S. patentapplication Ser. No. 10/836,740, filed Apr. 29, 2004, entitled“Compressed Porous Materials Suitable For Implant”, which is assigned tothe same assignee as this invention, and which disclosure is fullyincorporated by reference herein.

BACKGROUND OF THE INVENTION

The present invention relates to surgical devices for stabilizing and/orfusing adjacent bone structures, and, more particularly, to surgicaldevices for stabilizing and/or fusing the spine and for implantationbetween the vertebrae, including the intradiscal space. Generally, thisinvention concerns internal fixation devices, particularly spinal fusionand related implants.

Spinal degenerative diseases (e.g., stenosis, disc disease, spondylosis,etc.) trauma, aging, or a herniated disc can cause compression in thespine thus applying pressure to the nerve roots and/or spinal cord. Thecompression produces progressive pain, loss of movement and sensation,and sometimes, permanent disability. Spinal fusion is among thestandards of care for surgical decompression and stabilization of thespine. Fusion, known also as arthrodesis, is accomplished by theformation of an osseous bridge between adjacent motion segments. Thegoals of spinal surgery include relieving spinal cord/nerve compression,promoting spinal fusion, increasing stability, maintaining spinalalignment, and restoring disc height. Ideally, reconstructive surgerywould result in total spinal fusion with an excellent clinical outcome.

For over 40 years, removal of the problematic disc and fusion of theadjacent vertebrae has been the common treatment for degenerativediseases. The classical surgical procedure is discectomy and interbodyfusion with an iliac crest autograft with or without internal fixation.A discectomy typically requires the removal of a portion or the entireintervertebral disc. Different types of grafts (e.g., autograft,allograft, or synthetic ceramics) are used to fill the disc space.

Unfortunately, the use of bone grafts presents several disadvantages.Autogenous bone, which contains matrix molecules and living cells suchas osteoblasts that facilitate fusion, is the ideal bone graft; however,postoperative pain is often greater at the harvest site than thesurgical site. Additionally, autografts removed from a patient may notyield a sufficient quantity of graft material. Harvesting bone is alsoassociated with high rates of harvest site morbidity and can increasethe risk of infection and blood loss. Alternatively, allografts obviatethe need for bone harvesting, but have inconsistent mechanicalproperties. Allografts can also transmit diseases or cause infections,and they have unpredictable and slow fusion rates. Autografts andallografts alone may not provide the stability required to withstandspinal loads and are subject to collapse or failure due to a lack ofstrength.

In the mid-1970's, Bagby found the clinical results of harvest sitemorbidity to be unacceptable. In U.S. Pat. No. 4,501,269, he describesthe “Bone or Bagby Basket” to eliminate bone graft harvesting andpromote bone fusion. Due to the drawbacks of traditional fusiontechniques, his initial invention was important and innovative, and ithas continually been improved in both design and material selection.These interbody fusion devices are designed to stabilize the vertebralbodies, hold osteogenic material, and promote early stabilization andfusion. The rigidity and structural design of the devices must be ableto support the axial loads in the spine. Commercially available spinalinterbody fusion devices are made of stainless steel, titanium alloy,carbon fiber, or allograft bone. Often, these devices have void spacesor perforations to allow bone ingrowth.

While carbon fiber and metal interbody fusion devices offer strengthadvantages, they have several disadvantages. Metal interbody fusiondevices are a permanent foreign body and are difficult to remove duringrevision surgery. Due to the difference in mechanical properties of boneand metal, the main concern of metal interbody fusion devices isstress-shielding, which may cause bone resorption or osteopenia.Although these devices have demonstrated an ability to facilitatefusion, a sufficient fusion is not always achieved between the bonegrafts housed within the cage and the vertebral endplates. Achieving acomplete bony union in the middle portion of the cage has beenparticularly problematic. Clinical fusion outcomes may be difficult toassess with metallic interbody fusion devices due to the artifacts andscattering during postoperative CT or MRI scans. Often a complete bonyunion cannot be seen, making fusion results unreliable. Carbon fibercages are radiolucent and have properties, such as modulus ofelasticity, similar to bone; however, they are also a permanent foreignbody. Long-term results with metal and carbon fiber interbody fusiondevices are unknown due to the relatively recent development of theimplants. Metal cages have been known to fatigue and will eventuallyfail if a solid bony fusion is not achieved. Over time, metal and carbonfiber cages may migrate or have significant subsidence into thevertebral bodies.

Gjunter (U.S. Pat. No. 5,986,169) describes a porous (i.e., 8 to 90%porosity) material made of a nickel-titanium alloy. The pores form anetwork of interconnected passageways that permit fluid migrationthrough the material. The material may be used for biomedical implantsor non-medical applications. Kaplan (U.S. Pat. No. 5,282,861) andZdeblick et al. (U.S. Pat. No. 6,613,091) discuss a similar porousmaterial made of a carbon-tantalum composite that could be used tocreate an implant device. The elasticity of the porous materials issimilar to live bony tissue; however, most of the disadvantagesassociated with carbon fiber and solid metal internal fixation devicesstill apply to the porous nickel-titanium and carbon-tantalum alloymaterials. For example, the porous metal implants remain permanentlyimplanted in the body.

To avoid the disadvantages of metal and carbon fibers devices,bioresorbable materials have been used for years as sutures, boneplates, screws, pins, and other medical devices. A few advantages ofbioresorbable implants include biocompatibility, predictabledegradation, and complete resorption via natural pathways by the bodyover a period of time. Polymers are advantageous over otherbioresorbable materials, such as ceramics, because they have hightoughness and are highly reproducible. The toughness significantlyreduces the danger of polymers failing by brittle fracture.Bioresorbable polymers can be formed into spacers, wedges, threadedcages, and a variety of other shapes (e.g., spinal interbody fusiondevices).

Bioresorbable implants are transparent to x-rays, and therefore allow,for example, postoperative clinical assessment of a bony union, therebyovercoming one disadvantage of metallic implants. They can perform allthe requirements of an interbody cage by providing immediate stability,maintaining foraminal distraction, restoring disc height, and allowingbone ingrowth and fusion. Bioresorbable interbody fusion devices can beproduced that provide sufficient strength retention (up to 12 months orlonger) in order to allow fusion to occur, then resorb after they are nolonger needed. They have the compressive strength to withstand and carrythe spinal axial loads; however, they have a modulus of elasticitysimilar to bone, which limits stress-shielding. Bioresorbable implantdevices may feature or contain osteogenic material to attract bone andcells to the implant. Additionally, the bioresorbable devices may behydrophilic and/or porous. Porous, hydrophilic devices promote themigration of fluid material into the implant, thus allowing a widevariety of tissue ingrowth. The porous bioresorbable implants are fullycapable of being replaced by the patient's own bone growth.

Lynch (U.S. Pat. No. 5,306,303), McKay (U.S. Pat. No. 6,346,123) andWebb (U.S. Pat. No. 6,503,279) all describe bioresorbable, porousceramic materials that may be used in medical implants. McKay and Webbspecifically describe an intervertebral fusion device. Due to thebrittle nature of ceramic materials, particularly as degradation occurs,the disclosed materials may not withstand the axial loads or cyclicloading of the implant site (e.g., spine) without fracture, collapse,and ultimately device failure.

McKay (U.S. Pat. Nos. 5,702,449 and 6,039,762) describes a spinal cagewith an inner core of porous biocompatible material, preferably porousceramic, which allows tissue ingrowth, and an outer body that canwithstand compressive loads. The porous biocompatible material mayprotrude from the outer shell to permit contact with the vertebralbodies. The implant design with the resorbable inner core does not allowfor the use of a bone graft within the device. A high strength outershell may provide sufficient support; however, it brings concomitantproperty mismatch with natural bone. Bioceramics as used to form theouter shell are brittle and may fracture under high spinal loads.

Moumene and Serhan (U.S. Pat. No. 6,569,201) disclose a fusion cage witha structural bioresorbable layer disposed upon the outer surface of anon-resorbable support. The purpose of the non-resorbable support is toact as a scaffold for the bioresorbable layer and to hold a bone graftor osteogenic material. The bioresorbable layer would resorb over time,gradually increasing the loading on the bone graft and fusion cage. Ifthe bioresorbable layer and bone graft degrade before fusion can occur,the non-resorbable support may cause stress-shielding. Depending on thethickness of the bioresorbable layer, complete degradation of the layermay cause a great decrease in disc space height. The non-resorbablesupport will remain a permanent foreign object in the body.

Gresser et al. (U.S. Pat. Nos. 6,241,771 and 6,419,945) describes aspinal interbody fusion device composed of 25-100% bioresorbablematerial. The device is composed of a resorbable polymer that canproduce acidic products upon degradation and includes a neutralizationcompound to decrease the rate of pH change as the device degrades. Inorder to withstand the maximum physiologic loading, of at least 10,000 N(the maximum expected lumbar load), the disclosed device must bereinforced with fibers. The device is not porous, consequently limitingbone ingrowth. Similar to metal interbody fusion devices, the device mayhave void spaces to hold osteogenic materials, such as bone grafts orother osteogenic material. The disclosed device will slowly degrade andlose strength over time with complete resorption predicted to occur byone year. Clinically, complete fusion and bony union may take longerthan one year in unstable patients. If fusion of the endplates throughthe disk space does not occur, the short-term resorption of the devicemay lead to collapse of the disk space.

Bioresorbable interbody spinal fusion devices offer solutions todisadvantanges related to bone grafts and metal and carbon fiber cages.Autografts require bone graft harvesting, which causes postoperativepain and morbidity at the harvest site. Allografts put the patient atrisk for infection or transmitted diseases. Metal and carbon fiber cagesremain permanent foreign bodies. Metal cages can cause stress-shieldingand make fusion assessment difficult. They may also migrate from theimplantation site or subside into the vertebral bodies. A need existsfor an interbody spinal fusion device that achieves a successful fusionand bony union while avoiding drawbacks associated with the use of metaland carbon fiber devices or bone grafts.

SUMMARY OF THE INVENTION

The present invention is a compressed porous matrix material forapplication to a tissue site in order to promote new tissue growth. Oneaspect of this invention is glass transitional deformation orcompression of a porous polymeric composition to create a high-strengthmaterial that retains the benefits imparted by its porous nature. Thecompression of the porous composition creates a three-dimensionalmulti-laminated structure having equivalent mechanical properties tosolid (monolithic) polymeric structures without the problems associatedwith such structures. Compression can affect and create a new structurefrom the non-compressed porous matrix material. Certain compressionmethods may create collapsed pore walls that form thin, overlappinglaminate walls. Because the laminate walls are formed from thinoverlapping laminate walls wherein the walls form a continuousintercommunicating network within the device, the laminate layers arethereby limited in the amount that they can slide, thus eliminating thissliding or delaminating mode of failure, which may otherwise be seen.Depending on the amount of compression, the porous matrix material mayhave a few collapsed pores or may be completely made up of thin,collapsed pores. Variations in the compression method can createcollapsed pores that did not form thin laminate walls, but instead thepores are condensed to a fraction of their original size. Due to poresthat collapse or give way first, the pores throughout the material mayvary in size. This compressed porous matrix material may be fabricatedinto many different devices for various applications in the body, whichwill be discussed later.

Any biocompatible polymeric material, which can to be fabricated into aporous matrix by those skilled in the art, is envisioned to bemanufactured by the methods disclosed herein. Methods for creating aporous structure are well known to those skilled in the art (e.g.,oil-water emulsions, lyophilization, precipitation, particulateleaching, critical gas blowing, gas forming polymerizations, etc.). Asan example, one method involves dissolving a polymer in a solvent (e.g.,acetone, chloroform, ethanol, dioxane, NMP, t-butanol, water, etc.) andfiltering. The material is then treated to remove the residual solvent.Precipitating the polymer, evaporative distillation, lyophilizing thesolution, or other methods may be used to remove the solvent, thusforming a porous polymeric material.

Another example involves dissolving a polymer in an organic solvent toprepare a polymer solution of high viscosity, or mixing a polymersolution in an organic solvent that does not dissolve the polymer toconcentrate the solution as a gelatinous precipitate. A salt ishomogeneously mixed with the polymer solution or gelatinous precipitateto give a polymer/salt/organic solvent mixed gel. The organic solvent isremoved from the mixed gel through techniques known in the art (e.g.,air dry, vacuum dry, sublimation, etc.) to produce an organicsolvent-free polymer/salt composite. The composite is submerged in anaqueous solution or acidic solution to cause the salt to leach out atroom temperature to form a porous three-dimensional polymeric structure.The porous three-dimensional polymer structure useful for the presentinvention may contain open celled intercommunicating pores and/or closedcelled non-communicating pores.

The resultant porous matrix material is compressed by force; preferably,at temperatures at or above the materials glass transition temperature,but below the melt temperature. Any method of compression known by thoseskilled in the art is conceivable for this invention, including, but notlimited to, using hydraulically or pneumatically powered platens orpistons to compress the porous matrix material. Other methods includeusing a screw or an arbor press to compress the material. Compression isdefined as a method for applying force to a porous matrix material inorder to alter the size, shape, mechanical/material properties, and/orstructure of the original material. The compression has many variables,including the amount of force/pressure used, the percent compression ofthe original height, the direction of compression, etc. The percentcompression directly corresponds to the amount of porosity aftercompression. It should be noted that although compression reduces theoverall porosity of the material, surface area of the pores is minimallyaffected. The compression temperature can also be varied to create thedesired properties of the material.

Those skilled in the art will recognize that polymers without a glasstransition temperature can still be utilized in creating theabove-described invention by means of inducing pseudo glass transitions.The simplest means of creating a pseudo glass transition is byincorporation of a plastisizer or plastisizing the polymer with smallamounts of solvent. Other methods include, but are not limited to,quenching and cycling the temperature just above and below the meltpoint of the polymer. One skilled in the art will also recognize thatthese methods for creating pseudo-glass transition may also beeffectively utilized for polymers having glass transition states.

Another aspect of this invention relates to controlled stretching andmolding of a porous matrix material. Heating to temperatures above theglass transition temperature allow the porous polymer to soften andcontract. If contraction is prevented and a force in a new direction isapplied, the malleable polymer can be stretch to the extent that theporosity can collapse allowing the porous matrix to be pressed into orover a mold. Cooling at this time will lock in the new shape. The areaof polymer that has been shaped is different than the unaltered areas.This is due to the forced alignment of the polymer partitions. Molds maybe tailored to impart anisotropic effects at discrete locationsthroughout the implant, through creating areas of higher flow (i.e.,more stretching) as well as areas of very low flow. Therefore,properties may be tailored by location and degree. Unlike thecompressing method described above, this method has the ability toincrease the surface areas within the porous matrix as the porosity isreduced.

Various medical uses of the above-described invention are describedbelow. Other features or advantages of the present invention will beapparent from the following drawings and detailed description of theinvention, as well as from the claims.

BRIEF DESCRIPTION OF THE FIGURES

FIGS. 1A and B illustrate the porous matrix material, including is porestructure, before (1A) and after (1B) being compressed.

FIG. 2A illustrates the porous matrix material between two compressivedevices.

FIG. 2B shows the porous matrix material being compressed by the topcompressive device.

FIG. 2C shows the porous matrix material being compressed by bothcompressive devices.

FIG. 2D shows the porous matrix material being compressed by bothcompressive devices in a heated atmosphere.

FIGS. 3A and B illustrate the three-dimensional compression of a porousmatrix sphere, including its pore structure before (3A) and after (3B)being compressed.

FIGS. 4 A and B illustrates the three-dimensional compression of aporous matrix cylinder, including its pore structure before (4A) andafter (4B) being compressed.

FIG. 5 shows the pore structure of compressed porous matrix materialthat contains various additive materials.

FIG. 6A is a perspective view of one embodiment of the implant.

FIG. 6B is a perspective view of one embodiment of the implantcontaining an osteogenic material.

FIG. 6C is a perspective view of the embodiment from FIG. 6A and one ofthe anatomical locations that is suitable for treatment by the implant.

FIG. 7 is a perspective view of an alternative embodiment of the implantand one of the anatomical locations that is suitable for treatment bythe implant.

FIG. 8A is a perspective side view of an alternative embodiment of theimplant.

FIG. 8B is a perspective front view of the embodiment from FIG. 8A.

FIG. 9A is a perspective view of an alternative embodiment of theimplant.

FIG. 9B is a perspective view of an alternative embodiment of theimplant.

FIG. 9C is a perspective view of an alternative embodiment of theimplant.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

The object of the invention is an implantable prosthesis, constructed ofa compressed porous polymeric material. The construction of theprosthesis is such that it is capable of absorbing energy and supportinglarge compressive loads utilizing less mass of material than would benecessary in the formation of a solid polymer prosthesis. Additionally,the device has advantages over metal prostheses, including theresorbable nature of the prosthesis and transient nature of its stressshielding.

While working to create a low porosity material, a new and unique methodto control or alter the porosity within a porous material wasdiscovered. In a preferred embodiment, the method for preparing thehigh-density porous matrix involves:

-   -   a) creating a high porosity polymeric matrix by methods known in        the art;    -   b) inducing glass-transition within said porous polymeric        matrix;    -   c) applying a compressive force within one or more dimensions to        achieve a new size or shape; and    -   d) cooling the porous polymer out of the glass-transition        wherein the polymer matrix maintains the new size or shape.

A method of producing a high density porous matrix that may experienceglass transition after being compressed to a new size or shape involves:

-   -   a) creating a high porosity polymeric matrix by methods known in        the art;    -   b) inducing glass-transition within said porous polymeric        matrix;    -   c) applying a compressive force within one or more dimensions to        achieve a new size or shape;    -   d) holding the porous polymer matrix above glass transition at        the new size and shape for a period of time allowing the        molecular chains within the matrix to rotate or move to a lower        energy state; and    -   e) cooling the porous polymer out of the glass-transition        wherein the polymer matrix maintains the new size or shape.

Those skilled in the art will recognize porous materials that are noteither brittle and/or susceptible to fracture (e.g. elastic polymers) donot need to be at glass transition prior to the compressive step,although, it may be required to place them in a state of glasstransition while maintaining them in a compressive state to lock thematerial into the new conformation. Additionally, glass-transition maynot be necessary for porous materials that do not naturally re-expand(e.g., porous metal) or that have been or will be contacted with asecond substance that serves as a binder (e.g. bio-glue, adhesive,polymer solution) to lock the porous polymeric matrix in the compressedstate. This binder can be an external coating or a substance that isflowed into the porosity and functions to hold the overlapping laminatewalls together post compression. This binder may also be a temporary(e.g. biodegradable, dissolvable, heat sensitive) material that allowsthe compressed porous material to re-expand at a later time. This can beuseful in filling voids that have small openings or for delivery of acompressed pellet through a cannula to a surgical site (e.g. spine)where it is allowed to re-expand and provide support. It would beobvious to one skilled in the art that a binder would not be necessaryfor a re-expanding foam if the compressed foam only re-expanded whenplaced in contact with body fluids or at body temp.

When the porous matrix material is compressed, some or all of the poresmay be sacrificed and collapse to form laminate walls. The pores arelimited in the amount they may move within the material structure beforethey must absorb the compression and/or torque. Some pores aresacrificed, giving way to other pores, which may stay structurallyintact giving the matrix material an unique toughness not seen in priorart materials. Depending on the method, direction, and amount ofcompression, the sacrificed pores could give way in different modes(e.g., collapsing, folding, slipping, reducing in overall size,narrowing, etc.). The movement of some pores within the material and thesacrificing of other pores may cause the material to compress, thuschanging the material and mechanical properties of the porous matrixmaterial. The collapsing of the pores will have a direct effect on theporosity of the porous matrix material. The porosity will decreaseduring compression as pores are sacrificed and relieve compressivestress. The term “sacrificed” is used to both describes the initialcollapse of pores during manufacture and any further changes to thepores in response to forces on the finished device. Toughness ispartially imparted by the ability of a localized area within a device toindependently accommodate stresses.

Compression can create a new structure within the porous matrixmaterial. Certain compression methods may create collapsed pore wallsthat form thin, overlapping laminate walls. The laminate walls may beadjacent to each other, as in the case where compression has beenapplied to completely collapse the pores. Alternatively, the laminatewalls may have some space or material between the walls, such that theyare not in direct contact with each other (i.e., not adjacent). Becausethe laminate walls are formed from the collapsed pores, the layers arelimited in the amount that they can slide, thus eliminating adelaminating or sliding mode of failure. Variations in the compressionmethod and parameters can create collapsed pores that did not form thinlaminate walls, but instead pores are condensed to a fraction of theiroriginal size. Because pores closest to the applied force typicallycollapse or give way first, the pores throughout the material may varyin size, creating an altered pore size distribution throughout thematerial. Given walls of equal thickness, larger pores are more likelyto collapse than smaller pores. This can be used to reduce overallvariation in pore size. Other methods for compression can producetubular pores that are narrowed, lengthened, and/or collapsed. Thetubular pores can span the length of the material or be interconnected.In all cases, compression parameters may be modified to produce materialsuitable for end use as a medical device.

The structure of the porous matrix material also depends on the amountof compressive force applied to the material. The amount of compressionmay change the porosity of the porous matrix material. The pore sizedistribution will also be affected by the amount of compression as theporous matrix material may be compressed so that only certain areas havecollapsed pores, or so that all of the pores are sacrificed andcollapse. The direction of compression in relationship to the originalstructure of the porous matrix material can also affect the structure ofthe compressed porous matrix material. For example, if the initialporous matrix material has long tubular columns, a force appliedcircumferentially to the material will collapse the diameter of thecolumns; whereas a force applied parallel to the axis of the columnswill shorten the column length.

Compression of the porous matrix material can be controlled to createvarious structural patterns within the material; likewise, themechanical properties of the material may be altered to meet specificrequirements. The amount of compression is directly related to themaximum compressive load of the material. The more the material iscompressed, the greater the maximum compressive load will be. If amedical device fabricated from the compressed material must withstandloading from more than one direction, the compressed material can becompressed three-dimensionally to increase the mechanical strength ofthe material in all directions. If the medical device is axially loaded,the compressed material may be compressed in one direction to optimizethe mechanical properties of the material in that direction.

Generally, solid non-elastic, non-porous polymeric materials (i.e.,polylactides, poly-dl-lactide, etc.) have good mechanical strength;however, they are brittle and will catastrophically fail under highcompressive loads. Compressed porous material exhibits more ductilityand toughness compared to the same non-porous polymeric material. Thecompressive porous material has the ability to absorb energy bysacrificing pores. As compression continues beyond the point when allthe pores have collapsed, the material may expand slightly andmicrocracks will occur along its surface, thus avoiding catastrophicfailure.

Preferably, porous polymeric materials (fibrous and/or non-fibrous) arecompressed for the present invention, although it is also envisionedthat porous metallic materials (fibrous and/or non-fibrous) may also becompressed. It should be noted that many of the benefits imparted topolymeric materials, including ductility and toughness would also beimparted to compressed porous metallic materials. Thus it is anotherobject of the invention to create improved, lightweight porous metallicimplants useful in orthopedic surgery (e.g., artificial hip implants,bone plates, femoral nails, screws, etc.).

The temperature of the porous matrix material (e.g., glass transitiontemperature) during compression can greatly affect the behavior of thefinal material. More specifically, the relationship between thecompression temperature and the material's glass transition temperatureplays a vital role in the properties of the final material.Glass-transition is defined as the state during which the moleculesmaking up the matrix are free to move and rotate in an effort to achievea lower energy state. At temperatures narrowly above the glasstransition temperature, and below the melting temperature, the moleculealignment will occur more slowly than would alignment at temperaturesfurther above glass transition, but still below melting temperature.Those skilled in the art will recognize that polymers with an extremelyhigh glass transition temperature, or even no glass transition, canstill be utilized in creation of the present invention by means ofinducing pseudo glass transitions. The simplest means of creating apseudo glass transition is by incorporation of a plasticizer orplasticizing the polymer with small amounts of solvent. This can be doneby blending a plasticizer into the polymer or exposing the polymer to anatmosphere of molecules that would solvate the polymer at higherconcentrations. Other methods include, but are not limited to, quenchingand cycling the temperature just above and below the melt point of thepolymer. Quenching allows crystalline polymers to become amorphous for ashort period of time and may in turn create a pseudo glass transitionbelow the melt point of the polymer. Cyclic heating and cooling of apolymer just above and below its melt point can be used to simulate aglass transition by retarding collapse of the porous structure. If thematerial is compressed below the glass transition temperature, stresscan be locked into the material. If the material is then exposed totemperature at or above the glass transition temperature, the stresswill be relieved and the porous matrix material may expand and possiblyreturn to its pre-compression dimensions. Yet, if the material iscompressed at a temperature at or above the glass transition temperatureor brought up to glass transition after compression while still beingcompressed, the polymer chains in the material are free to rotate andassume a lower energy state. This may eliminate the stress in thecompressed material and the material will retain its dimensions evenwhen exposed to temperature at or above the polymer's glass transitiontemperature for a period of time.

If not compressed initially into the final shape, after being compressedand removed from the compression device, the porous matrix material maybe machined into a new shape or design with various features. Machiningprocesses for polymeric materials are well known to those skilled in theart. (e.g., coring, milling, sawing, lathing, etc.) As an example, atubular device could be machined by coring out the inner diameter andthen using a lathe to create the proper outer diameter.

The porous matrix material may be compression molded into an initial orfinal design of a medical device. If the device has complicatedgeometry, various features may be machined after compression molding,creating a refined shape for the device. As discussed above, thematerial and mechanical properties of the final device can be altered bythe compression or mold temperature, the amount of overall compression,the design of the mold, etc. The porous matrix material may becompressed before molding or all the compression may occur during themolding process. The direction of compression before or duringcompression molding may also affect the mechanical properties of thedevice. For example, a cylinder of porous material may bethree-dimensionally compressed and then compression molded into athreaded bone screw. Additionally, if the mold is heated and compressionis performed rapidly, only those areas in direct contact with the moldwill achieve glass transition, and collapse in response to compression.In this manner, a device having bi-modal pore structure can be created,as the pores in the center remain unaltered by compression.

The prosthesis may be sterilized by any method known in the art (e.g.exposure to ethylene oxide, hydrogen peroxide gas plasma, e-beamirradiation or gamma irradiation, etc.). The sterilization processminimizes the opportunity of infection to occur as a result of theimplant.

In a preferred embodiment of the invention, a porous medical device ismanufactured from a resorbable material, although this is not meant toexclude the use of non-resorbable polymers and metals. Differentpolymers, molecular weights, additives, processing methods, andsterilization methods can be used to control the resorption rates ofresorbable polymers. Resorption rates can be adjusted to be shorter forapplications that require mechanical strength for only a short period oftime or longer for applications that require mechanical strength to bepresent for a longer duration. The materials of the construct may befibrous or non-fibrous. Examples of resorbable polymers that can be usedto form medical devices are shown in Table 1. These materials are onlyrepresentative of the materials and combinations of materials that whichcan be used as implant materials.

TABLE 1 Examples of Bioresorbable Polymers for Construction of theMaterial of the Current Invention. Alginate Aliphatic polyestersCellulose Chitin Chitosan Copolymers of glycolide Copolymers of lactideElastin Fibrin Glycolide/1-lactide copolymers (PGA/PLLA)Glycolide/trimethylene carbonate copolymers (PGA/TMC) GlycosaminoglycansLactide/tetramethylglycolide copolymers Lactide/trimethylene carbonatecopolymers Lactide/e-caprolactone copolymers Lactide/s-valerolactonecopolymers L-lactide/dl-lactide copolymers Methyl methacrylate-N-vinylpyrrolidone copolymers Modified proteins Nylon-2 PHBA/g-hydroxyvaleratecopolymers (PHBA/HVA) PLA/polyethylene oxide copolymers PLA-polyethyleneoxide (PELA) Poly (amino acids) Poly (trimethylene carbonates) Polyhydroxyalkanoate polymers (PHA) Poly(alklyene oxalates) Poly(butylenediglycolate) Poly(hydroxy butyrate) (PHB) Poly(n-vinyl pyrrolidone)Poly(ortho esters) Polyalkyl-2-cyanoacrylates PolyanhydridesPolycyanoacrylates Polydepsipeptides Collagen Types 1 to 20 Nativefibrous Soluble Reconstituted fibrous Recombinant derivedPolydihydropyrans Poly-dl-lactide (PDLLA) Polyesteramides Polyesters ofoxalic acid Polyglycolide (PGA) Polyiminocarbonates Polylactides (PLA)Poly-1-lactide (PLLA) Polyorthoesters Poly-p-dioxanone (PDO)Polypeptides Polyphosphazenes Polysaccharides Polyurethanes (PU)Polyvinyl alcohol (PVA) Poly-b-hydroxypropionate (PHPA)Poly-b-hydroxybutyrate (PBA) Poly-s-valerolactone Poly-b-alkanoic acidsPoly-b-malic acid (PMLA) Poly-e-caprolactone (PCL) Pseudo-Poly(AminoAcids) Starch Trimethylene carbonate (TMC) Tyrosine based polymers

For the purposes of promoting an understanding of the principles of thisinvention, reference will now be made to the embodiments illustrated inthe drawings, where like numbers refer to like components, and specificlanguage will be used to describe the embodiments and elements of theembodiments. It must be understood that no limitation of the scope orapplications of the invention is thereby intended. For ease ofunderstanding, pores are represented in the drawings by closed circles,it is recognized the pores may in fact be formed in various shapes,textures and interconnectivity (e.g., they may be interconnected orseparate, open cell or closed cell, organized or random, and/orreticulated structures).

Referring now to the drawings, FIG. 1A depicts the porous matrixmaterial 100 before any compressive force has been applied. The porousmatrix material 100 includes a large percentage of void space, which isoccupied by pores 110. The pores 110 form the structure within thepolymeric material 120. After being compressed, as depicted in FIG. 1B,the compressed porous matrix material 105 contains the same amount ofpolymeric material 125; however, the sacrificed, collapsed pores 115have reduced the porosity of the material.

In another embodiment, as depicted in FIG. 2A, uncompressed porousmatrix material 200 is placed between two devices capable of applyingcompressive force 210 (e.g., platens, pistons, etc.), which may or maynot be heated. The pores 220 and polymer material 230 define thestructure of the uncompressed porous matrix material.

As depicted in FIG. 2B, the compressive device 215 is actuated to createpartially compressed porous matrix material 240. Upon compression of thematerial 240, a gradient is formed, wherein the compressed pores 250first begin to collapse, while the pores 220 (depicted here in FIG. 2Bas the lower part of the material) furthest removed from the actuatedcompressive device 215 retain their original structure. This can beemployed to create an implant for biphasic tissues such as bone. Theportion containing the collapsed pores 250 resembling cortical bone andthe remaining portion remaining uncompressed 220 resembling cancellousbone.

As shown in FIG. 2C, dual gradient porous matrix materials 260 can beformed by compressing the porous matrix material with a plurality ofactuated compressive devices 215, actuated either in succession orsimultaneously. The compressed surfaces, containing the pores closest tothe actuated compressive devices 215, will contain the highestproportion of sacrificed or compressed pores 250. The next layercontains the partially compressed pores 280 which have started tocollapse, but initially will decrease in size before completelycollapsing or being sacrificed. The porous matrix material furthestremoved from the actuated compressive devices 215, in the middle of thedual gradient material 260, will have pores 220 that most closelymaintain their original structure and size. This multiple compressiontechnique depicted in FIG. 2C may be employed to create an implant for atriphasic tissue such as the skull, requiring an implant that mimics thetransitions from cortical bone (more solid) to cancellous bone (porous)and back to cortical bone.

As shown in FIG. 2D, an evenly and significantly compressed porousmatrix material may be created, such as by actuating the compressivedevices 215 acting upon porous matrix material, by completely collapsingand sacrificing every pore. As a result, the evenly and significantlycompressed material could be formed without any of the gradients createdin devices of FIGS. 2B and 2C. As seen in FIG. 2D, the sacrificed orcollapsed pores 250 can be distributed through out the material 290.This is useful in the creation of a superior implant to replace thosecurrently manufactured from cortical bone, metal, or solid polymers.

An evenly compressed porous matrix material 290 may also be created byactuating the compressive devices 215 upon the material, while it isexposed to a heated atmosphere (e.g., convection oven, environmentalcontrol chamber, etc.). The heated environment may be above the glasstransition temperature of the polymeric material. As a result, an evenlycompressed material 290 could be formed without being significantlycompressed and without any of the gradients created in the devices ofFIGS. 2B and 2C. As seen in FIG. 2D, the sacrificed or collapsed pores250 will be evenly distributed through the material 290.

It is envisioned that desired percentages of porosity or desired poreshapes and sizes can be created based on the amount and method ofcompression. Specific pore shapes (e.g., spherical, thin flat sheet,tubular, etc.) or sizes may promote different types of tissue ingrowth(e.g. bone or vascular tissue ingrowth). Based on desired porosity orpore structure, the porous matrix material may act as a cellularscaffold for various uses in tissue engineering.

In another embodiment, surfaces of the porous matrix material (whetherpartially compressed 240 depicted in FIG. 2B, a dual gradient material260 in FIG. 2C, or evenly compressed 290 as shown in FIG. 2D) while incontact with actuated compressive devices 215, which may or may not beheated, could have compressed pores 250 forming extremely thin sheets.The extremely thin compressed pores 250 may form laminate walls, thusproviding a confining matrix for confining new tissue growth within thedevice. This can be important for applications involving areas such asthe spine where vital neural and vascular tissues are exposed andvulnerable.

In another embodiment, as illustrated in FIG. 3A, an uncompressed shape,(e.g., a sphere) 300 of porous matrix material is to be subjected tocompressive forces in three dimensions, with the compressive forces tobe applied depicted by arrows 310. This three dimensional compressionmay be applied in a variety of forms, for example mechanical means ofcompression, or alternatively by exposing the sphere 300 to a highpressure environment (e.g., increased atmospheric or hydrodynamicpressure). Pores 320 within a polymeric material 330 create theuncompressed sphere's 300 structure. As depicted in FIG. 3B, afterapplication of compressive forces, the porosity and size of thecompressed sphere 340 have been decreased. Unlike two-dimensionalcompression, the pores 350 have not collapsed into thin, laminate walls.The three-dimensional compression resulted in compressed pores 350, byreducing the pores in size, rather than inducing collapse. This decreasein the size of the sphere may be caused by folding of the poresresulting in a decrease in the constrained area within each pore, or anincrease in wall thickness between the pores of the polymeric material(not shown). This embodiment may be implanted within the body forvarious purposes, for example as a device to promote staged delivery ofbiologically active agents or alternatively, the device or a section ofthe device may be used to create an implant in order to repair, replaceor supplement a body part (e.g., a chin or a cheek). The embodiment of athree dimensionally compressed shape may also be used to create a cellbased implant wherein the cells supported in the non-compressed centerof the device are protected from the body's immune system by thecollapsed porous exterior. This would be particularly useful insupporting and protecting transplanted tissue (autograft or xenograft)such as islate cells capable of producing insulin. While immune cellswould be prevented from entering the sphere 340 and destroying theislate cells, oxygen and nutrients would readily pass through thecollapsed pores 350. In turn, waste product and insulin would pass outof the sphere.

Two-dimensional compression may also be applied upon a shape (e.g., acylinder) as illustrated in FIGS. 4A and 4B. Like the sphere 300 of FIG.3A, the uncompressed cylinder 400 of FIG. 4A is composed of pores 410within a polymeric material 420. Two-dimensional compression may beapplied to the cylinder 400 by applying force around the circumferenceof the cylinder 400 while restricting elongation of its height. Thistype of two-dimensional compression may result in the smaller diametercompressed cylinder 430 of FIG. 4B. The compressed cylinder 430 mayfeature pores 440 that have been forced to narrow under two-dimensionalcompression yet maintain their relative height. If the elongation of thecompressed cylinder 430 is encouraged (e.g., by tension applied at oneor both ends of the cylinder), the pores within may narrow and lengthen(not shown). Depending on the amount of compression applied, the pores440 could form thin tubes running parallel to each other throughout theheight of the cylinder 430. Devices like this would be useful in variousmedical applications (e.g., as orthopedic rods, nerve guides, etc.).

It is recognized that the pores 440 can be compressed by a drawing orlengthening action of the cylinder 400. As porous materials are broughtabove glass transition, they soften and contract. If contraction isprevented and a force in a new direction is applied, the now malleablematerial may stretch to the extent that the porosity can collapse andthe void volume is lost. This will allow the porous material to beshaped by being compressed into, stretched into, or drawn over a mold.In this way, porous sheet material can be stretched into concave moldsor over convex molds allowing the formation of unique cup or cavityshaped sheets. In essence, the porous sheet material at or above glasstransition can be thermoformed by any method known to those skilled inthe art, including, for example, male/female molding and vacuum drawing.The area of the porous polymer that has been shaped is stiffer than theunaltered areas of the sheet. This is believed to be due to the forcedalignment of the polymer partitions defining the pores.

The forced alignment of the pores can also be used to create apseudo-elastic memory in non-elastic polymers. If a porous sheet isbrought above glass transition and drawn in a single direction, thepores can collapse in the transverse direction while elongating in thelongitudinal direction. After cooling below glass transitiontemperature, the sheet resists forces applied in the longitudinaldirection, but will easily expand in the transverse direction byallowing the elongated collapsed pores to open up as the entire sheetshortens in the longitudinal direction. If the force in the transversedirection is released, the sheet springs back to its elongated form.

This process can also be applied to the compressed cylinder 430 in FIG.4B. If the cylinder is compressed around its circumference with tensionapplied to both ends while being heated, the pores will be forced intoalignment while being narrowed and lengthened. After cooling down,tension could be applied at various locations around the center of thecylinder. As the cylinder expands and bows in the middle, the centralpores are widened, yet the top and bottom pores move closer to eachother. When the tension is released, the cylinder and pores return totheir normal compressed shape and size.

A device having elongated pores capable of widening movement in thetransverse direction could be used a ligament or tendon. In a tubularform, it could be useful as a vessel, nerve guide, esophagus or othertubular organs. Additionally, it could be used as a sleeve, sack, or bagstretched over or around implants (e.g., rods, nails, etc.) or used tohold materials, for example granular materials such as ceramics (e.g.,hydroxyapatite, tricalcium phosphate, etc.), or other materials such astissues (e.g., cells, bone chips, demineralized bone, bone marrowaspirate, etc.).

In another embodiment as illustrated in FIG. 5, a compressed polymermatrix material 500 may be created in a common shape (e.g., a block, asphere, etc.) and/or shaped, machined, or molded to fit a particularapplication, with the material further containing or coated with atleast one additive component 530. These additives may be associated withonly the surface 510 of the polymer matrix material, rather thanextending into the interior of the shaped material (e.g., serving as acoating or shell). Alternatively, the additives 530 may be distributedthroughout and incorporated into the material 500 and/or the pores 520,either in a random or non-random dispersion. In an embodiment of thedevice having a random dispersion of the additives 530, the additivesmay be uniformly distributed throughout the volume of the polymer matrixmaterial 500. In another embodiment, the additive 530 may be distributednon-randomly, i.e., having a non-uniform distribution of additive 530within the polymer material 500, or within a depot within the material500. The non-uniform distribution may impart a desired quality to thematerial (e.g., by selectively affecting a portion of the material 500,by providing the ability to deliver a drug or multiple biologicallyactive agents as a burst and/or over an extended period of time, etc.).In another embodiment, the additives 530 may be associated with only thepores 520 within the polymer matrix material 500. In any of theembodiments containing additives, the pores may be open or closed cell,random or interconnected.

In one embodiment, at least one of the additives 530 of FIG. 5 may serveto reinforce the polymer matrix material 500. The reinforcing additives530 serve to enhance the characteristics of the device, such asmechanical strength (e.g., modulus of elasticity, compressive strength,tensile strength, etc.) and biodurability (e.g., hydrolytic degradation,strength retention, etc.). This may be accomplished by incorporatingreinforcing elements (e.g., mesh, fibers, threads, screen, etc.) ontothe surface, or incorporated into the material 500 (e.g., uniformlydispersed or individual layers) and/or the pores 520 of the polymermaterial. To further improve the mechanical properties of the material,the reinforcing elements may be interwoven, layered, or compactedtogether during the manufacture of the uncompressed polymeric material,or as a result of compression in making the compressed polymericmaterial 500.

In another embodiment, at least one of the additives 530 of FIG. 5 mayinclude or be a biologically active agent (e.g., growth factors,demineralized bone material, cells, drugs, viruses, etc.). The uniqueporous structure of the compressed material 500 can be used to controlthe location and delivery of the biologically active agents. Theformation of the construct controls the flow of fluid (e.g., blood,interstitial, etc.) within the device allowing for tailored releaseproperties. The biologically active agents may be incorporated into thedevice along with reinforcing agents, in which case, it is recognizedthe biologically active agents may be mechanically or chemicallyattached or bonded to the reinforcing materials. Alternatively, it isalso recognized that any of the additives 530 (e.g., reinforcing orbiologically active agents) may be delivered together in the material500 of the device without being mechanically or biologically bonded.Examples of biologically active agents that may be delivered in thedevice are shown in following Table 2.

TABLE 2 Examples of Biological Active Ingredients Adenovirus with orwithout genetic material Alcohol Amino Acids L-Arginine Angiogenicagents Angiotensin Converting Enzyme Inhibitors (ACE inhibitors)Angiotensin II antagonists Anti-angiogenic agents AntiarrhythmicsAnti-bacterial agents Antibiotics Erythromycin PenicillinAnti-coagulants Heparin Anti-growth factors Anti-inflammatory agentsDexamethasone Aspirin Hydrocortisone Antioxidants Anti-platelet agentsForskolin GP IIb-IIIa inhibitors eptifibatide Anti-proliferation agentsRho Kinase Inhibitors (+)-trans-4-(1-aminoethyl)-1-(4-pyridylcarbamoyl)cyclohexane Anti-rejection agents Rapamycin Anti-restenosis agentsAdenosine A_(2A) receptor agonists Antisense Antispasm agents LidocaineNitroglycerin Nicarpidine Anti-thrombogenic agents ArgatrobanFondaparinux Hirudin GP IIb/IIIa inhibitors Anti-viral drugsArteriogenesis agents acidic fibroblast growth factor (aFGF) angiogeninangiotropin basic fibroblast growth factor (bFGF) Bone morphogenicproteins (BMP) epidermal growth factor (EGF) fibringranulocyte-macrophage colony stimulating factor (GM-CSF) hepatocytegrowth factor (HGF) HIF-1 insulin growth factor-1 (IGF-1) interleukin-8(IL-8) MAC-1 nicotinamide platelet-derived endothelial cell growthfactor (PD-ECGF) platelet-derived growth factor (PDGF) transforminggrowth factors alpha & beta (TGF-.alpha., TGF-beta.) tumor necrosisfactor alpha (TNF-.alpha.) vascular endothelial growth factor (VEGF)vascular permeability factor (VPF) Bacteria Beta blocker Blood clottingfactor Bone morphogenic proteins (BMP) Calcium channel blockersCarcinogens Cells Cellular materials Adipose cells Blood cells Bonemarrow Cells with altered receptors or binding sites Endothelial CellsEpithelial cells Fibroblasts Genetically altered cells GlycoproteinsGrowth factors Lipids Liposomes Macrophages Mesenchymal stem cellsProgenitor cells Reticulocytes Skeletal muscle cells Smooth muscle cellsStem cells Vesicles Chemotherapeutic agents Ceramide Taxol CisplatinCholesterol reducers Chondroitin Collagen Inhibitors Colony stimulatingfactors Coumadin Cytokines prostaglandins Dentin Etretinate Geneticmaterial Glucosamine Glycosaminoglycans GP IIb/IIIa inhibitors L-703,081Granulocyte-macrophage colony stimulating factor (GM-CSF) Growth factorantagonists or inhibitors Growth factors Bone morphogenic proteins(BMPs) Core binding factor A Endothelial Cell Growth Factor (ECGF)Epidermal growth factor (EGF) Fibroblast Growth Factors (FGF) Hepatocytegrowth factor (HGF) Insulin-like Growth Factors (e.g. IGF-I) Nervegrowth factor (NGF) Platelet Derived Growth Factor (PDGF) RecombinantNGF (rhNGF) Tissue necrosis factor (TNF) Transforming growth factorsalpha (TGF-alpha) Transforming growth factors beta (TGF-beta) VascularEndothelial Growth Factor (VEGF) Vascular permeability factor (UPF)Acidic fibroblast growth factor (aFGF) Basic fibroblast growth factor(bFGF) Epidermal growth factor (EGF) Hepatocyte growth factor (HGF)Insulin growth factor-1 (IGF-1) Platelet-derived endothelial cell growthfactor (PD-ECGF) Tumor necrosis factor alpha (TNF-.alpha.) Growthhormones Heparin sulfate proteoglycan HMC-CoA reductase inhibitors(statins) Hormones Erythropoietin Immoxidal Immunosuppressant agentsinflammatory mediator Insulin Interleukins Interlukin-8 (IL-8)Interlukins Lipid lowering agents Lipo-proteins Low-molecular weightheparin Lymphocites Lysine MAC-1 Methylation inhibitors MorphogensNitric oxide (NO) Nucleotides Peptides Polyphenol PR39 ProteinsProstaglandins Proteoglycans Perlecan Radioactive materials Iodine - 125Iodine - 131 Iridium - 192 Palladium 103 Radio-pharmaceuticals SecondaryMessengers Ceramide Somatomedins Statins Stem Cells Steroids ThrombinThrombin inhibitor Thrombolytics Ticlid Tyrosine kinase Inhibitors ST638AG-17 Vasodilators Histamine Forskolin Nitroglycerin Vitamins E C YeastZiyphi fructus

The inclusion of groups and subgroups in the tables is exemplary and forconvenience only. The grouping does not indicate a preferred use orlimitation on use of any material therein. For example, in Table 2, thegroupings are for reference only and not meant to be limiting in any way(e.g., it is recognized that the Taxol formulations are used forchemotherapeutic applications as well as for anti-restenotic coatings).Additionally, the table is not exhaustive, as many other drugs and druggroups are contemplated for use in the current embodiments. There arenaturally occurring and synthesized forms of many therapies, bothexisting and under development, and the table is meant to include bothforms.

In another embodiment, at least one of the additives 530 of FIG. 5 maybe in the form of particulate components or filler materials (e.g.,tricalcium phosphate, biphasic calcium phosphate, hydroxylapatite,calcium sulfate, tetracalcium phosphate, autologous bone graft,allograft bone matrix, polymers, microspheres, etc.), which enhance thefunctionality of the device. The particulate components may be deliveredwithin the polymeric material 500 in various forms (e.g., granules,chips, powders, gels, etc.). The incorporation of particulate componentsinto the polymeric material 500 may enhance the ability of the device toexhibit desirable biological qualities (e.g., cellular growth promotion,bioactive osteoconductivity, tissue ingrowth promotion, etc.).Furthermore, the particulate components may also serve to enhance themechanical strength of the material. A non-exhaustive list of additivematerials 530 that may be incorporated in the present invention in theform of particulate or filler materials is provided in Table 3.

TABLE 3 Examples of particulate or filler materials suitable for use inthe present invention Alginate Bioglass Calcium Calcium PhosphatesMonobasic Dibasic Tribasic Ceramics Chitosan Cyanoacrylate CollagenDacron Demineralized bone Elastin Fibrin Gelatin Glass Gold Hyaluronicacid Hydrogels Hydroxy apatite Hydroxyethyl methacrylate NitinolOxidized regenerated cellulose Phosphate glasses Polyethylene glycolPolyester Polysaccharides Polyvinyl alcohol Radiopacifiers SaltsSilicone Silk Steel (e.g. Stainless Steel) Synthetic polymers ThrombinTitanium

In another embodiment, at least one of the additives 530 of FIG. 5 mayserve to impart or create a microstructure within the macrostructure ofthe polymeric material 500. Preferably, the macrostructure may serve tomaintain the mechanical, architectural, and structural stability of thedevice and provide a biologically inert surface for tissue ingrowth. Themicrostructure additive may, in a preferred embodiment, serve to attractand nourish inbound cellular growth. The additive material 530 suitablefor creating a microstructure can be selectively varied within certainregions of the macrostructure to promote or deter different biologiccharacteristics critical to different tissue requirements. Themicrostructure creating additive 530 could be contained by orconcentrated within the compressed pores 520. The microstructure can bestrategically located within one or more compressed pores 520. Themicrostructure creating additive 530 may also be on the surface 510 ofthe macrostructure. When located within collapsed intercommunicatingpores, the microstructure may prevent complete collapse of the pore. Thevolume of microstructure can be used to control the percentage collapseof each pore. The space created by the microstructure, as well as thehydrophilic/hydrophobic properties of the microstructure, influence therate at which fluids and/or cells flow into and out of the collapsedpores. In this way, microstructure could be used to control the releasekinetics of other additive materials, such as biologically activeagents, supported within the polymer or within the microstructureitself, from the device.

It is recognized that any of the above-described additive agents 530 maybe used alone or in combination with other additive materials. It isalso recognized that individual components making up the additivematerials may serve a dual purpose as an additive (e.g., acting as abiologically active agent and a reinforcing agent concurrently). Whenmore than one additive 530 is used within the polymer material 500, theadditives may function separately, or have a synergistic effect, whereinthe activity of one class of additive 530 helps the activity of theother class of additive component 530. The additives may physically bebonded together, or merely placed in proximity with each other, or evendistributed randomly or non-randomly without any interrelationship. Itis also recognized that based on the physical characteristics of theadditive components, some of the components may not resorb or may resorbinto the body at a different rate from other components, or have similaror different temporal qualities, such that the effects of the differentadditives may persist for various durations.

In another embodiment, shown in FIG. 6A, a resorbable spinal implant inthe form of an interbody fusion device (e.g., spinal cage, spacer,wedge, etc.) 610 may be created from the compressed porous matrixmaterial 600. An interbody fusion device 610, once implanted usingtechniques known in the art, may serve to restore the disc space in aspinal column. An interbody fusion device created utilizing the material600 of the present invention may be used to provide a large surface areato provide for adequate bone ingrowth, thereby eliminating the need forthe prior art technique of bone harvesting for autografts to be used increating a spinal implant.

The device of the present invention may also be constructed as a spinalimplant for posterolateral fusion (not shown). A posterolateral spinalimplant spans and contacts the transverse processes of adjacentvertebrae. The posterolateral implant made in accordance with thepresent invention will maintain a space above and across the transverseprocesses and facilitate new bone formation.

The device of the present invention may be constructed as an anteriorfusion spinal implant (not shown). An anterior spinal implant wouldfasten to two vertebrae and span the operative disc space, therebyserving to restrict motion and promote fusion through bone growth.

In an embodiment of the present invention, the compressed pores 620within spinal implant device 610 may be of any size or shape andarranged in any orientation suitable for use as a spinal implant.Preferably, the compressed pores 620 would be formed as thin, laminatesheets, which enable the device 610 to withstand both large compressiveloads and cyclic loading. In a more preferred embodiment, the structureand design of the device 610 will give it desirable mechanicalproperties (e.g., compressive strength, modulus of elasticity, tensilestrength, etc.) similar to cortical and/or cancellous bone.

In another embodiment, one or more channels 630 (e.g., pores, holes,slots, perforations, etc.) may be molded, machined, or drilled into thematerial of the present invention, for example as shown in FIG. 6Adepicting a spinal implant. Channels 630 can be created in anyorientation or direction into or through the device 610. Channels 630may pass completely through the device 610 thereby forming at least onevoid or reservoir from top to bottom or from side to side, at any angle.The size of the channels 630 can vary or may be the same. It isrecognized the compressed pores 620 and channels 630 may provide astructural function and/or biological function. In the example of aspinal implant, the channels 630 may provide a scaffold forvascularization and/or bone ingrowth, in order to facilitate theoccurrence of spinal fusion. The channels 630 may also serve tofacilitate resorption of the polymer from which the device has been madeby reducing the bulk or amount of polymer per device. Additionally, lesspolymer per device may lead to decreased manufacturing costs, as rawmaterial consumption is reduced. For example, for similar sized objects,one solid polymer and another being 10% porous material prepared asdescribed herein, the porous material utilized 10% less raw material,and may possess markedly better physical characteristics.

With reference to FIG. 6B, a channel (e.g., osteoconductive pore) orchannels created in the material of the present invention may also beuseful for the introduction of various biodegradable materials ormatrices 640. For example, material constructed as a spinal fusionimplant device 610 may feature a channel or hole (as depicted by channel630 of FIG. 6A) that has been filled with material 640. In a preferredembodiment, one material 640 that may be contained in the channel(s) isosteogenic grafting material (e.g., bone grafts, demineralized bone,bone void fillers, hydroxypatite, bone chips, bioceramics, etc.) topromote bone ingrowth into and through the device 610. The channel 630may be packed with the osteogenic material, which may be provided invarious forms (e.g., chips, strips, sheets, sponges, gels, etc.Potential biodegradable matrices 640, which may at least partially fillthe channel(s), may include beneficial materials (e.g., collagen sponge,collagen-ceramic composites, open-cell polylactic acid (OPLA), etc.) Thematerials or matrices may act as carriers for bone growth factors orosteogenic proteins, such as naturally or genetically engineered bonemorphogenic proteins (i.e., BMP-2, BMP-4, etc.).

Referring again to FIG. 6A, in one embodiment of the device a channel630 may also be used to accommodate a suitable tool (not shown) tofacilitate insertion of device 610 into the living being. For example,in the case of a spinal implant, a tool may be inserted into the channel630, thereby allowing controlled placement of the spinal implant into avertebral disc space. The channel 630 and corresponding tool may or maynot be threaded, or provide some temporary locking arrangement (e.g.,keyed, friction fit, etc.) to provide extra control during implantation,wherein movement of the tool relative to the channel 630 may be limited.

In various embodiments, the compressed porous matrix material 600 can beconveniently machined or molded during compression to form spinalimplants with complex geometries and various features. For example,polymer spinal implants may be created in a variety of differentconfigurations (e.g., a horizontal threaded cylinder, a vertical ring,an open box cage, etc.). A gripping means 650 may be provided to ensureadequate stability of the implanted spinal device 610. The grippingmeans may be any features that prevent the device from sliding orundesirable shifting from the implantation site. These gripping means650 may operate as a friction fit or incorporate locking elements (e.g.,teeth, serrations, ridges, grooves, threads, wedges, blocks, pins,nails, screws, staples, etc.), which may be machined or molded into thedevice 610. For the example of a spinal fusion implant, the grippingmeans 650 have the ability to grasp the vertebral endplates and resistlateral movement, thus helping to prevent the implant from migrating outof the vertebral disc space. Additionally, the gripping means 650 mayserve to impart increased surface area to the implant device 610, inorder to allow the device to withstand spinal pressures. Any recessescreated in the gripping means 650 (e.g., the spacing between consecutiveteeth or threads) may also serve to facilitate bone ingrowth that mayaid in anchoring the device in place. In an embodiment relying onthreads functioning as the gripping means 650, the threads may bemachined or molded on the outer surfaces of a compressed porous matrixshaped material (e.g., a dowel) to form a device similar to a threadedscrew. The threads allow easy and controlled insertion into thevertebral disc space.

In another embodiment, a device 610 may be shaped like a rod (notshown). The rod may feature a gripping means (e.g., ridges or teeth). Itis recognized that a device in the shape of a rod may beneficiallyincorporate a taper, such that one end is larger than the other, oralternatively, the rod may lack a taper.

In another embodiment, as illustrated in FIG. 6A, a spinal cage 610 canbe fabricated from the porous matrix material 600 into a spacer in theshape of a wedge. The wedge shaped device 610 may serve to providevertebral spacing and may aid in interbody fusion between vertebrae. Thecage or spacer device 610 may be tapered to provide the correctorientation to the vertebrae with which the device is in contact and canalso serve to keep the device in place. It is recognized the spacer maybe machined into any other shape or size (e.g., cylindrical, as shown inFIG. 6A, rectangular, kidney shaped, etc.) in order to conform to theshape of the vertebral endplates. Gripping means 650 may be machinedinto the cage 610 for additional spinal stability. The gripping means650 may be any height, shape, or size, depending upon the intended useof the device 610. The gripping means 650 may be located on one or moresurfaces of the device 610 and oriented in one or more directions on thedevice 610. As shown in FIG. 6C, the device 610 is sized and configuredfor engagement between two vertebrae 660. Preferably, the implant device610 has a height approximately equal to or slightly greater than theheight of the intervertebral disc space 670.

In various embodiments, the porous matrix material may be composed oflayers of the same or different types of polymers. Two or more differentporous polymer may be included in one device. It is recognized that thisinvention may be useful for medical devices that require specificabilities, material or mechanical properties, or biological conditionsto function optimally in the body. For example, devices may undergochanges in loading over time, require specific degradation rates, may beloaded differently across the surface of the implant, etc. In order toaccommodate the special requirements of some devices, in an embodiment,two or more different compressed porous matrix materials may be layered(e.g., stacked on one another, or alternatively side-by-side) to formthe device. Alternatively, the same porous matrix material may becompressed under different conditions. In these layered embodiments, thelayers of compressed material or materials may possess variable materialand structural characteristics (e.g., degradation rates, flexibility,drug delivery rates, etc.). The layers may or may not be fused together.The layers may be compressed by different methods or by differentamounts. The layers may provide the device the ability to bemulti-functional. For example, it is recognized that one or more layerscan perform one function (e.g. provide structurally integrity, maintainshape, etc.) for the device while one or more other layers performanother function (e.g., drug delivery, allow bone ingrowth, etc.).

In another embodiment, the compressed porous matrix material can bemachined or molded into any configuration, such as an internal fixationdevice for use in surgical repair, replacement, or reconstruction ofdamaged bone in any area of the body. Internal fixation devices may besuccessfully employed for many conditions and applications (e.g.,orthopedic, spinal, maxiofacial, craniofacial, etc.).

Another possible embodiment of the invention is an internal fixationdevice, as shown in FIG. 7, where a plate 700, is affixed in ananatomical location 710. A plate 700 may be machined or molded withfixation holes 720 that will allow fixation to bone by various meansknown in the art (e.g., staples, screws, tacks, etc.). During thesurgical implantation procedure, the fixation holes 720 may also becreated to fit the anatomical location 710. Holes 720 createdcontemporaneously with implantation of the plate may allow more accurateplacement or fitting of the plate 700, consequently a more effectiveapplication of the invention. The plate may be useful as a graftcontainment device for the repair or reconstruction of defects, such asthose caused by surgery, tumors, trauma, implant revisions, infections,and also for joint fusion.

In another embodiment, illustrated in FIGS. 8A and 8B, compressed porousmatrix material may be machined or molded into an interbody fusionplating system 810, which is a device that is a combination of a cage orspacer 820 and a plate 830. The interbody fusion plating system 810provides the benefits of a resorbable cage or spacer 820, and furtherincorporates the advantages of a plate 830. The plate 830 may increasefusion rates by acting as an anterior tension band, reducing motion andmovement at the implantation level. The plate 830 will prevent themigration and loosening of the cage 820. The resorption of the plate 830over time will gradually increase loading on the cage 820 and bonytissue, promoting fusion. The interbody fusion plating system 810 may befabricated as one solid device or two single devices that can beconnected and used together or used separately. The plate 830 may bemachined with fixation holes 840 that will allow fixation to bone bymeans known to those skilled in the art (e.g., staples, screws, tacks,etc.). Alternatively, the fixation holes 840 may be created in thedevice contemporaneously with implantation, in order to ensure properplacement of the fixation holes in the device.

Various representative embodiments are illustrated in FIGS. 9A, 9B, and9C, wherein medical devices may be fabricated into any configurationfrom the compressed porous matrix material. Such devices may be used inany field wherein the functionality of the porous polymer material as afixation device may be useful, including but not limited to the fieldsof internal fixation, trauma repair, sports medicine, etc. For example,these devices suitable for bone and soft tissue fixation may includescrews 900, rods, pins 910, tacks, arrows, staples 920, washers, nails,anchors, etc. These devices may be used in many applications requiringfixation devices, such as the repair of fractured bones.

The following examples are given for purposes of illustration to aid inunderstanding the invention and it is to be understood that theinvention is not restricted to the particular conditions, proportion,and/or methods set forth therein.

Example 1

The objective of this example is to compare the physical properties ofdifferent Poly-l-lactide (PLA) porous matrix materials after beingcompressed 0, 40, 60, and 80% of its original height. Static axialcompression tests were performed to measure the maximum compressiveloads of the porous matrix materials after being compressed to differentpercentages of their original height. The compression tests willdemonstrate the compressed material's mechanical properties can bealtered and controlled over a wide range of possible values. The finalproperties of the compressed material are determined by the propertiesof the starting material and the amount of compression used. The finalproduct is a material that has tensile and compressive strengths similarto that of non-porous polymer yet is not as stiff or subject to failureby cracking as non-porous polymer. The mechanical and porosity testswill assure a device fabricated from compressed porous matrix material(e.g., a spinal interbody fusion device) is able maintain its porosityand absorb fluids, while still being able to withstand large stressesand loads it may be subject to (e.g., the maximum physiologic loadingexpected in the lumbar spine of at least 10,000 N).

The compression test procedure for the compressed porous matrixmaterials are based on ASTM standards D1621-94, Standard Test Method forCompressive Properties of Rigid Cellular Plastics, and D1667-97,Standard Specification for Flexible Cellular Materials—Vinyl ChloridePolymers and Copolymers (Closed-Cell Foam). The only polymer used forthis example was Poly-l-lactide (PLA). The porous matrix materials wereproduced by methods known to those skilled in the art. The porous matrixmaterials can be created with porosities that initially range from 98%to 86% or lower. At least five cylindrical specimens (15 mm in diameterand 15 mm in height) were machined from each material. An axial load wasapplied via a materials testing system to each cylindrical specimen at arate of 12.5 mm/min until a stopping point of 50% strain. The loadversus displacement curves were recorded. For each test, the maximumcompressive load and compressive modulus of elasticity were calculatedand recorded.

Additional material property tests included porosity and wettability.The wettability and porosity were measured to determine the effects ofcompression on the porous material. The porosity of each material wasmeasured before and after being compressed using a Helium Pycnometer,which determines the density and volume of a sample by measuring thepressure change of helium in a calibrated volume. The wettability(ability of the material to absorb fluids) of the material wasdetermined on a pass/fail basis after compression, subjectivelyassessing the ability of the porous material to absorb fluids.

The different PLA materials with initial porosities ranging from 97% to86% were compressed by 0% to 84% of its original height. Up to themaximum compression of 84%, the materials maintained their wettabilityand a percentage of the original pre-compressed porosity. The material'sstrength after compression was directly related to the initial porosityand amount of compression. For example, the initial porosity andcompression strength of the uncompressed materials ranged from 97%porosity with 30 N of compressive strength to 86% porosity with 624 N ofcompressive strength. At 40% compression, the strength and new porosityfor the two materials with the lowest and highest initial porositiesranged from 67 N and 94% porosity (97% initial porosity) to 1348 N and75% porosity (86% initial porosity). The compressive strength andporosity ranged from 101 N and 90% (97% initial porosity) to 2249 N and71% porosity (86% initial porosity) after 60% compression. The finalcompression set point of 80% resulted in compressive strengths andporosities ranging from 326 N and 84% porosity (97% initial porosity) to4889 N and 57% porosity (86% initial porosity).

In order to find a compressive strength greater than 10,000 N, thematerial with the lowest initial porosity (86%) was compressed by 84% ofits original height. At the 84% compression, the maximum compressiveload was 12,985 N and the actual measured porosity was 41%. Relying onthe following equation, where theoretical porosity can be calculated as1−[(1−initial % porosity)/(1−% compression)], the theoreticallycalculated porosity would have been around 13%, with the differencebetween the theoretical and actual porosity percentage values mostlikely being due to the sample not being restrained as compression wasapplied, and allowed to expand horizontally beyond the 15 mm diameter ofthe original sample. Had there been some restraint against expansionwhile being compressed, the percentage porosity would have been reducedto less than 14% porosity, down from the original 86% of the initialmaterial. The maximum compressive load of 12,985 N is above the maximumexpected physiological spinal loading of 10,000 N. The porous matrixmaterial can be produced with a lower initial porosity and compressed byvarious methods (previously described) to increase the maximumcompressive strength, providing a significant safety factor compared toboth typical and maximum physiological spinal loading.

The results from the porosity, wettability, and compression tests provethat PLA porous matrix material can be compressed by various degrees togive a wide range of compressive strengths while still maintaining itsporosity. By altering PLA porous material and the amount of compression,any amount of porosity and compressive strength may be created. Thecompressive strengths were found to range from 30 to almost 13,000 N.The compressed material may be useful as an internal fixation device,such as a spinal fusion cage. A spinal fusion cage made of compressedporous matrix material would be able to withstand the maximumphysiologic loading expected in the lumbar spine of at least 10,000 N.The maximum compressive load found in this example of 12,985 N is abovethe maximum physiological spinal loading. Even accounting for thehorizontal expansion during compression, the increased area that resultsis less than the surface area of a lumbar vertebra and thus stillexceeds the expected load. The porous matrix material can be producedwith a lower initial porosity and compressed by various methods(previously described) to increase the maximum compressive loadproviding a significant safety factor compared to both typical andmaximum physiological spinal loading.

Example 2

While Example 1 demonstrated that PLA porous matrix materials could becompressed, this example serves to illustrate that porous matrixmaterials made of different polymers can also be compressed and willcompare the physical properties of the two compressed materials.Polylactide/Poly ε-Caprolactone (PLA/PCL) andPoly(desaminotyrosyl-tyrosine ethyl carbonate) (PDTE) Carbonate wereused to create two different porous matrix materials. The compression,porosity, and wettability tests described in Example 1 were used to testthese materials.

Static axial compression, wettability, and porosity tests were conductedas described in the EXAMPLE 1.

Before compression, the porosities of the PLA/PCL and PDTE Carbonatewere 92% and 94%, respectively. Up to 40% compression, the materialsshow little to no change in porosity. At 80% compression, the morebrittle porous material (PDTE Carbonate) had a porosity of 73% comparedto 66% porosity for the PLA/PCL material. It should be noted that, as inEXAMPLE 1, the samples were not restrained from expanding horizontallyduring compression; therefore the actual measured porosity values areslightly different from theoretically calculated porosity values.

The maximum compressive strength results showed significant differencesin the mechanical strength of the compressed materials. At 80%compression, the PDTE Carbonate had a maximum compressive strengthgreater than 1500 N compared to a compressive strength of 450 N for thePLA/PCL material. At 87% compression, the PLA/PCL with 59% porosity wasable to withstand a maximum compressive load of 581 N.

The objectives of this study were to determine if different materials(other than PLA) could be compressed and to compare the material andmechanical properties of two different porous matrix materials (PLA/PCLand PDTE Carbonate) after being compressed different percentages oftheir original height. Due to the elasticity of the PLA/PCL material, itwould only hold its shape if compressed at temperatures near its glasstransition temperature. The PDTE Carbonate could be compressed with orwithout heat and hold its compressed shape. After compression, eachmaterial still retained a high percent of its porosity and was able toabsorb fluids. The compressive strength results from the compressivetests were significantly different for each material. The PLA/PCLmaterial had a compressive strength much less than the PDTE Carbonatematerial. The elasticity of the PLA/PCL material prevents it from beinga material able to withstand large compressive loads. This study provesthat it is possible to compress elastic and brittle materials, as wellas non-lactide materials.

Thus since the invention disclosed herein may be embodied in otherspecific forms without departing from the spirit or generalcharacteristics thereof, some of which forms have been indicated, theembodiments described herein are to be considered in all respectsillustrative and not restrictive, by applying current or futureknowledge. The scope of the invention is to be indicated by the appendedclaims, rather than by the foregoing description, and all changes whichcome within the meaning and range of equivalency of the claims areintended to be embraced therein.

1. A porous implantable device suitable for implantation in a livingbeing, said implantable device comprising a high density porous polymermaterial having at least one property that is characteristic of at leastone of cortical bone and cancellous bone, said at least one propertybeing selected from the group consisting of compressive strength,modulus of elasticity, and tensile strength, wherein said high densityis obtained by a process comprising compression applied to at least oneporous polymer material, wherein said compression process causes thesacrifice of at least some pores in said porous polymer material,thereby forming a new shape, and further wherein said compression takesplace in at least one of (i) a temperature above the glass transitiontemperature for said polymer, and (ii) when said polymer is in aplasticized condition, whereby molecules of said polymer move and rotateto achieve a lower energy state, thereby causing said high densityporous material to permanently lock in the new shape, and wherein saiddevice is arranged for implantation into said living being.
 2. Theporous implantable device of claim 1, wherein said at least one polymermaterial is resorbable.
 3. The porous implantable device of claim 1,wherein said pores are sacrificed to create laminar walls.
 4. The porousimplantable device of claim 1, wherein said high density porous materialis further machined to a final shape.
 5. The porous implantable deviceof claim 1, further comprising at least one additive component.
 6. Theporous implantable device of claim 5, wherein said additive component isdistributed uniformly throughout the implantable device.
 7. The porousimplantable device of claim 5, wherein said additive component isdistributed on outside surfaces of the implantable device.
 8. The porousimplantable device of claim 5, wherein said additive component isdistributed within said pores of said implantable device.
 9. The porousimplantable device of claim 5, wherein said additive component isdistributed in a portion of said implantable device.
 10. The porousimplantable device of claim 5, wherein said additive component isarranged to create a microstructure.
 11. The porous implantable deviceof claim 10, wherein said microstructure is arranged to encourage tissueingrowth.
 12. A high density porous material comprising at least somesacrificed pores, suitable for implantation into a living being, havingat least one mechanical property being at least characteristic of atleast one of cortical bone and cancellous bone, and being obtained bythe process comprising the steps of: a) providing at least one highporosity material comprising a plurality of pores in the form of cells,wherein said pores are defined by pore walls; b) inducing aglass-transition state within said at least one high porosity material;c) applying a compressive force within one or more dimensions while saidhigh porosity material is in said glass transition state, saidcompressive force being sufficient to cause said molecules of said highporosity material to move and rotate to achieve a lower energy state,thereby causing said high porosity material to achieve a new size orshape, and further to sacrifice at least a portion of said pores,thereby creating said high density porous material; and d) cooling saidhigh density porous material out of said glass transition state, whereinsaid high density porous material permanently maintains the new shape orsize.
 13. The high density porous implantable material of claim 12,wherein said high porosity material further comprises at least onepolymer material.
 14. The high density porous implantable material ofclaim 13, wherein said at least one polymer material is resorbable. 15.The high density porous implantable material of claim 12, wherein saidprocess further comprises the step of: e) machining said high densityporous material to a shape.
 16. The high density porous implantablematerial of claim 12, wherein said high porosity material furthercomprises at least one additive component.
 17. The high density porousimplantable material of claim 16, wherein said additive component isdistributed uniformly throughout said high porosity material.
 18. Thehigh density porous implantable material of claim 16, wherein saidadditive component is distributed on outside surfaces of said highporosity material.
 19. The high density porous implantable material ofclaim 16, wherein said additive component is distributed within saidpores of said high porosity material.
 20. The high density porousimplantable material of claim 16, wherein said additive component isdistributed in a portion of said high porosity material.
 21. The highdensity porous implantable material of claim 16, wherein said additivecomponent is arranged to create a microstructure.
 22. The high densityporous implantable material of claim 21, wherein said microstructure isarranged to encourage tissue ingrowth.
 23. A porous implantable devicecomprising a porous polymer that is compressed by at least about 40percent and no more than about 87 percent of its original height, isarranged to maintain its compressed shape, and has at least onemechanical property being at least characteristic of bone.
 24. Theporous implantable device of claim 12, wherein said at least onemechanical property is selected from the group consisting of compressivestrength, modulus of elasticity, and tensile strength.
 25. The porousimplantable device of claim 23, wherein said at least one mechanicalproperty is selected from the group consisting of compressive strength,modulus of elasticity, and tensile strength.
 26. The porous implantabledevice of claim 23, wherein said bone comprises at least one of corticalbone and cancellous bone.